Radiation detectors are particularly used in computed tomography (CT) scanners and will be described with particular reference thereto. However, the invention also finds use in DF (diffraction) and RF (radio frequency) imaging, X-ray fluoroscopy, radiography, and other examination systems for medical and non-medical examinations.
Computed tomography (CT) imaging typically employs an X-ray source that generates a beam of X-rays traversing an examination area. A subject arranged in the examination area interacts with and absorbs a portion of the traversing X-rays. A two-dimensional radiation detector including an array of detector elements is arranged opposite the X-ray source to detect and measure intensities of the transmitted X-rays.
Typically, the X-ray source and the radiation detector are mounted at opposite sides of a gantry which rotates so as to obtain an angular range of projection views of the subjects. In some configurations, the X-ray source is mounted on the rotating gantry, whereas the radiation detector is mounted on a stationary gantry. In either configuration, the projection views are reconstructed by using filtered back-projection or another reconstruction method to produce a three-dimensional image representation of the subject or of a selected portion thereof.
The radiation detector may include an imaging plate consisting of an array of imaging elements, such as scintillation crystals, which produce bursts of light, called scintillation events, in response to X-rays. Such radiation detectors may also include an array of photodetectors such as a photodiode array which is arranged to view the scintillation crystals and produce analog electric signals indicative of the spatial location and intensity of the scintillation events. Imaging plates, for use in CT scanners and general medical examinations, include an assembly of pixels being independently responsive to the incident X-rays and generating electric signals, which are used to generate a digital image. In some detectors, the scintillator assembly includes an array of individual crystals which are assembled together or cut from a common scintillator plate, e.g. by dicing or other semiconductor manufacturing techniques.
Most CT manufacturers today make X-ray detector arrays, wherein each detector comprises one or more scintillators and one or more photodiodes. The X-ray detectors comprise blocks of crystalline or ceramic X-ray scintillator material which emit light, separated from each other by white spacers or separators and being glued to the front surface of silicon photodiode arrays. The white separators or spacers, which are made of light-reflecting material, usually comprise an epoxy resin selected for radiation hardness, with a titanium dioxide filler to make it white. The function of the light-reflecting material is to reflect light, generated by scintillation when X-rays are absorbed in the body of the scintillator, downwardly into the sensitive region of the photo-detecting element, to avoid loss upwardly, or scattering sideways into neighbouring dixels (detector pixels).
The detector array may have many or even hundreds of detector pixels, or dixels, and is optically coupled to and juxtaposed upon a matching silicon photodiode array. The silicon photodiode array collects the light emitted by the scintillators and generates electric charges that are electronically processed and used to display voxel characteristics in the subsequent CT image.
However, problems exist in the known X-ray detector arrays. As the X-ray detector arrays grow in size, thermal expansion problems between the stiff epoxy resin having a high coefficient of thermal expansion (CTE) and the fragile silicon chip having a low CTE may cause delamination, especially when the assembly suffers from extreme temperatures. This may happen during delivery of the scanner to a hospital in winter, when there may be temperatures below −20° C.
Moreover, white separators or spacers made by means of this known technology must be fairly thick. The efficiency of a white reflecting layer at wavelength λ is defined by the scattering coefficient Sλof Kubelka and Munk, which is related to the layer thickness d and the diffuse reflectance Rλ, as defined by their well-known formula
      s    λ    =            R      λ              1      -                        R          λ                ⁢        d            Typically, scattering coefficients not much larger than 2000 cm−1 can be achieved by using epoxy resins whose refractive index generally exceeds 1.5. This means that a separator having a thickness of 100 μm will transmit 5% of the light as crosstalk. This is particularly important when it is desired to reduce dixel size to improve the spatial resolution of the CT image.
Furthermore, white coatings at the outside edge of an array, where space for coatings is limited, must also be relatively thick. A coating having a thickness of 50 μm will lose 9% of the light incident upon it.
Accordingly, the invention preferably seeks to mitigate, alleviate or eliminate one or more of the above-mentioned disadvantages singly or in any combination.